1. Field of the Invention
The present invention relates generally to imaging devices for nuclear medicine, and more specifically relates to gamma ray or scintillation cameras and methods of correcting images obtained by such cameras for distortions due to nonlinear spatial response of the camera to incident radiation events.
2. Background and Prior Art
In nuclear imaging, a patient is injected with or swallows a radioactive isotope which has an affinity for a particular organ. Gamma rays are then emitted from the organ of interest, and are detected by a gamma ray or scintillation camera device which forms an image of the organ based on the concentration and distribution of the radioactive isotope within the organ.
The scintillation camera as utilized in nuclear medicine is a well known device. The original scintillation camera or "Anger camera" (named after the inventor) is described in U.S. Pat. No. 3,011,057. The Anger camera uses a scintillation crystal, such as a NaI crystal, which absorbs incident gamma rays from the object under study and interacts with the gamma ray to produce light events. An array of photomultiplier tubes is placed adjacent to the crystal in order to detect and amplify these light events so as calculate the spatial location and energy level of the incident gamma ray to produce a two dimensional image of the object which then may be displayed on a CRT or printed as a hard copy.
For each incident gamma ray which interacts with the crystal, the electronic circuitry of the camera combines the output of the photomultiplier tubes and individually computes the spatial (x,y) coordinates and the energy signal (z) for the detected radiation event. An image is generated by plotting the x,y position of a large number (typically millions) of such events. Since cosmic gamma rays and gamma rays which have been scattered will be incident on the detector in addition to incident gamma rays emitted from the radioactive isotope within the organ of interest, the energy signal z is used to identify certain detected gamma rays as being desired for contribution to the image, from among all gamma rays which are incident on the detector. The energy signal z has a functional relationship to the total energy of an incident gamma ray and thus incoming gamma rays are discriminated on the basis of the amplitude of this energy signal.
It is well known in the art that images formed from calculated coordinates of light events detected on the crystal of the camera head contain certain distortions. One such distortion is a spatial dislocation or linearity distortion, in which the coordinates of an event which occurs at a certain location on the crystal will be computed as being at a location which diverges from the true location. This distortion occurs because the calculated position of an incident gamma ray is dependent upon and varies with the location of the light event in the crystal with respect to the photomultiplier tubes. The effect of this is that light events in certain regions of the detector are moved toward the centers of the photomultiplier tubes and thus increasing the apparent density of events, leaving other regions where the perceived density is lower. This creates a so-called "barrel and pincushion" effect which causes an image to have a very nonuniform appearance.
Corrections for such spatial dislocation errors is known in the art and is disclosed in U.S. Pat. No. 4,212,061 to Knoll et al., which is incorporated by reference herein. Spatial coordinate correction factors are typically calculated by presenting a radiation image of a known pattern such as a grid of parallel lines or points, to the gamma camera detector. The spatial coordinates computed by the detector circuitry are then compared with the known coordinates of the grid pattern in order to calculate spatial coordinate correction factors which function to move the computed event positions to the actual, true positions. The calculation of spatial coordinate correction factors is conventionally performed with a gamma camera by utilizing the radioisotope of most common usage. In nuclear medicine, this is typically .sup.99m Tc which emits a gamma ray of 140 keV. Additionally, cobalt 57 which emits a gamma ray of 122 keV may be used.
In nuclear medical imaging applications, it is sometimes desired to simultaneously image gamma rays from radioisotopes which emit gammas at more than one energy, or to image gamma rays from two different radioisotopes. Different energy gamma rays produce different numbers of light photons when they are absorbed in the scintillation crystal. The number of light photons produced when a photoelectric absorption of a gamma ray occurs in the crystal is roughly linearly proportional to the energy of the absorbed gamma ray. For example, a gamma ray of 200 keV would produce approximately twice the number of scintillation photons as a gamma ray of 100 keV, assuming each was absorbed in the identical location in the scintillation crystal, and that each was totally absorbed by the photoelectric effect. Small differences occur in the calculated position of events of differing energies due to nonlinearities in the electronic circuitry processing the signals from the photomultiplier tubes generated through absorption of the scintillation photons.
As such, the nonlinearities associated with the spatial coordinate computation of light events are not only dependent on the spatial position of the event in the scintillation crystal, but are also dependent on the energy of the incident gamma ray. Therefore, linearity correction factors obtained for gamma rays of one particular energy will not be identical to correction factors obtained for radioisotopes emitting gamma rays at a different energy absorbed in the same location.